Dual inlet mixed-flow blood pump

ABSTRACT

A mixed-flow blood pump presents features of both axial-flow and radial-flow pumps. This mixed-flow blood pump comprises a stationary housing structure defining at least one blood inlet, a blood outlet, and a blood conduit between the at least one blood inlet and the blood outlet, and a rotative impeller mounted in the blood conduit. The at least one blood inlet, the blood outlet, the blood conduit and the rotative impeller have respective structures and configurations that operate the mixed-flow blood pump at a given point of a maximum hydraulic efficiency curve relating a specific pump rotational speed and a specific pump diameter. This given point is located within a transition region of the maximum hydraulic efficiency curve between axial-flow and radial-flow pumps.

FIELD OF THE INVENTION

The present invention relates to a mixed-flow blood pump displayingcharacteristics of both radial-flow and axial-flow pumps.

BACKGROUND OF THE INVENTION

The present specification mentions a number of references which areherein incorporated by reference.

In North America, heart related diseases are still the leading cause ofdeath. Among the causes of heart mortality are congestive heart failure,cardiomyopathy and cardiogenic shock. The incidence of congestive heartfailure increases dramatically for people over 45 years of age. Inaddition, a large part of the population in North America is nowentering this age group. Thus, the people who will need treatment forthese types of diseases comprise a larger segment of the population.Many complications related to congestive heart failure, including death,could be avoided and many years added to these persons' lives if propertreatments were available.

The types of treatment available for patients of heart failure depend onthe extent and severity of the illness. Many patients can be cured withrest and drug therapy but there are still severe cases that requirevarious heart surgery, including heart transplantation. Actually, themortality rate for patients with cardiomyopathy who receive drug therapyis about 25% within two years and there still is some form of thesediseases that cannot be treated medically. One of the last options thatremain for these patients is heart transplantation. Unfortunately,according to the procurement agency UNOS (United Network for OrganSharing in the United States), the waiting list for hearttransplantation grows at a rate of more than twice the number of heartdonors.

Considering these facts, it appears imperative to offer alternativetreatments to heart transplantation. The treatment should not only addto a recipient's longevity but also improve his quality of life. In thiscontext, mechanical circulatory support through Ventricular AssistDevices (VAD) is a worthwhile alternative given the large deficiency inthe number of available organ donors. In the 1980's, successfulexperiments with mechanical hearts and VADs serving as a bridge totransplantation increased significantly. The accumulated knowledge inall aspects of patient care, device designs and related problems led tothe use of VADs as permanent implants. Now, it appears appropriate toaddress the problem of end stage heart failure with permanent mechanicalheart implants. Among the various mechanical support devices, axial-flowVADs with a projected life span of five to ten years provides a veryinteresting approach. It is estimated that eight thousand (8,000)patients per year in Canada and seventy-six thousand patients (76,000)per year in the United States could benefit from VADs.

In 1980, the National Heart, Lung and Blood Institute (NHLBI) of theUnited States defined the characteristics for an implantable VAD(Altieri, F. D. and Watson, J. T., 1987, “Implantable Ventricular AssistSystems”, Artif Organs, Vol. 11, pp. 237-246). These characteristicsinclude medical requirements including restoration of hemodynamicfunction (pressure and cardiac index) avoidance of hemolysis, preventionof clot formation, infection and bleeding, and minimisation of theanti-coagulation requirement. Further technical requirements include:small size, control mode, long life span (>2 years), low heating, noiseand vibration.

VADs can be used in several circumstances where a patient has poorhemodynamic functions (low cardiac output, low ejection fraction, lowsystolic pressure). Whatever the origin of the cardiac failure, the goalof the VAD is to help the heart in its pumping action. The VAD reducesthe load on the heart by enhancing circulation, thus restoring thehemodynamic functions and providing improved end organ perfusion. Manycurrent VAD devices can achieve these goals; however they are notoptimal, and hemolysis and thrombus formation are still importantproblems requiring investigation.

In the 1970's, the first approach to the problem of mechanical supportwas to imitate as much as possible the heart physiology. This resultedin the development of several pulsatile devices, some of these initialdesigns being still in use. The first developments were pneumaticallydriven devices while a second generation of pumps was electricallyactuated. In the 1990's, a new generation of pumps has emerged whichaddresses certain problems associated with previous devices (size andpower consumption). These pumps are non-pulsatile devices divided intotwo main categories: radial-flow blood pumps, and axial-flow bloodpumps.

In a non-pulsatile VAD, an impeller is enclosed in a housing andcontinuously rotates to produce a pumping action. The faster therotation, the higher the blood flow. These devices are callednon-pulsatile or continuous because they provide for a constant bloodflow. Most axial-flow blood pumps operate around 10 000 RPM (RotationsPer Minute). However, in in-vivo conditions, there is a dynamic range(about 1000 RPM around the operating point) over which the output flowis pulsatile. Since the native heart is still contracting, a pressuredifference between the ventricle (inlet) and the aorta (outlet) iscreated. This pressure variation will produce a variation in the pumpflow. The range of rotational speed over which pulsatile flow occurs issmall; at lower speed back flow is observed (in diastole) and at higherspeed the load on the heart is reduced to zero. In the latter case, nopressure variation occurs resulting in non-pulsatile flow.

Many advantages are associated with the use of non-pulsatile VADs andthey all have a strong impact on the physiology as well as on theclinical management. These advantages include:

Size:

-   -   Non-pulsatile VADs provide a much smaller volume than pulsatile        VADs, around 25 cc for an axial-flow blood pump, and 100 cc for        a radial-flow blood pump, compared to 150 cc and more for        pulsatile devices. For the sake of comparison a complete        axial-flow VAD is usually smaller than the graft used for        pulsatile pump. The clinical impact is the possibility to use        this type of VADs in small adults as well as in children. Also,        the small dimensions allow for placement of the pump in a more        orthotopic position; that is, in the thorax near the heart        instead of the upper abdomen. This eliminates the use of long        cannula passing through the diaphragm. Furthermore, for        axial-flow VADs, the shape and size can be selected allowing for        placement of the VAD in an intra-ventricular position.

Power:

-   -   The electrical power required to drive a non-pulsatile VAD is        lower than for pulsatile VADs.

Simplicity:

-   -   Non-pulsatile VADs are mechanically simpler than pulsatile VADs;        they do not require complex structures such as valves,        diaphragms, blood sacs, vents or compliance chambers.        Non-pulsatile continuous VADs are made of a simple motor to        which is coupled an impeller contained in a housing. One        important advantage of a simple mechanical design is its        extended durability. Durability as long as five to ten years        (Nosé, Y., 1995a, “Can We Develop a Totally Implantable Rotary        Blood Pump?”, Artificial Organs, Vol. 19, pp. 561-562; and        Jarvik, R. K., 1995, “System Consideration Favoring Rotary        Artificial Hearts with Blood-Immersed Bearing”, Artificial        Organs, Vol. 19, pp. 565-570) could be achieved with continuous        VADs compared with two years for pulsatile VADs (Pierce, W. S.,        Sapirstein, J. S. and Pae, W. E., 1996, “Total Artificial Heart:        From Bridge to Transplant to Permanent Use”, Ann Thorac Surg,        Vol. 61, pp. 342-346). In principle, this would allow not only        to use a non-pulsatile VAD as a bridge to transplantation but        also as long term mechanical support.

Hemolysis:

-   -   Hemolysis, or tearing of the red blood cells, can be estimated        in vitro with a parameter called the Normalised Index of        Hemolysis (NIH).

Infection:

-   -   Interestingly, the probability of infection is reduced with        continuous VADs. This is due in large part to the transcutaneous        vent of a pulsatile VAD which is an open door for opportunist        infections and therefore requires daily cleaning.

Patient Issues:

-   -   Non-pulsatile VADs require less maintenance allowing the patient        a greater autonomy. Also, most patients with a VAD are        discharged from the hospital and returned to a normal life after        about a month. Presently, because of the vent in pulsatile VADs,        patients cannot take a bath or swim since water could enter the        motor compartment. Continuous VADs are less restrictive and        allow the patient to practice more activities.

Radial-flow blood pumps were first used in cardio-pulmonary bypass forheart surgery. Based on results obtained with the Bio-Medicus pump(Medtronic Bio-Medicus Inc., Eden Prarie, Minn.), several groups decidedto develop much smaller radial-flow blood pumps so that they could betotally implantable. In radial-flow blood pumps, the rotation of theimpeller produces a centrifugal force that drags blood from the inletport on top to the outlet port at the bottom. To produce rotation of theimpeller, the impeller is coupled to an electric motor. This coupling ismade either (a) magnetically by means of permanent magnets located underthe impeller and on the rotor of the motor or (b) mechanically by meansof a shaft interposed between the impeller and the motor's rotor.Magnetically coupled devices generally show better functionality becauseno seal is required between the motor and the impeller shaft.

A problem related to radial-flow blood pumps is that although they aremuch smaller than pulsatile pumps, they are still too large to betotally implanted in a human thorax thus eliminating anyintra-ventricular implantation.

To overcome the above-mentioned problem related to radial-flow bloodpumps, axial-flow ventricular assist blood pumps were developed. Theseaxial-flow blood pumps can decrease the hemolysis rate by decreasing thetime of exposure of the blood to friction forces and by reducing theintensity of these forces. Another interesting advantage is thataxial-flow blood pumps are generally much smaller than radial-flow bloodpumps.

The first commercially available axial-flow blood pump was the Hemopump™(Medtronic Inc. Minneapolis, Minn.) used as short term circulatorysupport. Based on the good results obtained with this pump, severalgroups have initiated the development of totally implantable axial-flowVADs for long term use.

A few axial-flow VADs are presently under intensive development.Examples are: the Jarvik 2000™, by the Texas Heart Institute (Houston,Tex.); MicroMed DeBakey™ by MicroMed Technology (also of Houston, Tex.);and HeartMate II™, by Thoratec (Pleasanton, Calif.). All of these groupshave already started in-vivo experiments on animals and humans althoughthey still perform in-vitro trials. An overview of the operation ofthese pumps is given hereinbelow.

Jarvik 2000™

-   -   This axial-flow blood pump comprises two stators, one at the        inflow and one at the outflow. These stators have two functions:        they support the bearing shaft around which the impeller will        rotate (middle part) and they modify the blood flow path. The        inflow stator initiates the rotation of the flow so that the        blade tip of the rotor does not create too much shear stress on        the blood cells. The outflow stator straightens the flow so that        blood from the pump enters the blood stream with a generally        axial-flow profile. Permanent magnets are enclosed in the center        of the impeller and two motor windings are located in the casing        on each side of the rotor. This configuration constitutes a DC        brushless motor; this is a simple and durable motor which        minimises the number of mechanical parts. The power cable is        connected directly to the DC brushless motor controller to        change motor speed.    -   To implant the device, the chest is opened by means of a left        thoracotomy and no cardiopulmonary bypass is used. The pump        axial outflow is anastomosed to the aorta with a Dacron™ graft.        Then a ventriculotomy is made to insert the pump into the        ventricle through a sewing ring attached to the apex.

Micromed DeBakey™

-   -   The Micromed DeBakey™ VAD is very similar to the Jarvik 2000™        design. Indeed, this concept of VAD is based on a DC brushless        motor with blood-immersed bearings, a central impeller and two        fixed side pieces. The Micromed VAD is described in U.S. Pat.        No. 5,527,159 (Bozeman, Jr. et al.) issued on Jun. 18, 1996.    -   Blood enters on the left side and passes through a flow        straightener preventing pre-rotation thereof. Then, the blood        reaches the inducer/impeller; the inducer initiates rotation of        blood before this blood reaches the impeller. However, it should        be noted that the impeller produces the effective pumping        action. Finally, a flow diffuser converts the tangential flow        into an axial flow. The inducer comprises three blades and the        impeller is provided with six blades. The three blades of the        inducer co-extend with three associated blades of the impeller.        Each blade of the impeller contains eight cylindrical permanent        magnets. Finally, a winding is placed outside the pump to        complete the motor assembly, and the rotor is supported by a        pair of bearings.

HeartMate II™

-   -   This pump is similar to the Micromed DeBakey™ and the Jarvik        2000™ pumps. It is placed next to the heart and is connected        between the apex of the heart and the aorta. It is also a        sealed-bearing type pump and, accordingly, requires a purge        system. This system has a second pump which injects 15 ml/day of        sterile solution in the sealed area. A pump without purge system        is now under development. Three animals have been supported for        one month with the HeartMate II™ (Konishi, H., Antaki, J. F. et        al., 1996b, “Long-term Animal Survival with an Implantable Axial        Flow Pump as a Left Ventricular Assist Device”, Artificial        Organs, vol. 20, pp. 124-127).

Other axial-flow blood pumps have been proposed. For example, U.S. Pat.No. 5,205,721 (Isaacson) issued on Apr. 27, 1993 discloses an axial-flowblood pump having a hydrodynamically suspended rotor centrallypositioned with respect to the stator. Hydrodynamic bearings are createdby two spaces in which blood must flow to create the hydrodynamicsupport; this in turn produces shearing forces applied to the blood. Inaddition, Isaacson teaches three types of impeller blades.

U.S. Pat. No. 5,211,546 granted to Isaacson et al. on May 18, 1993teaches an axial-flow blood pump which is similar to that of U.S. Pat.No. 5,205,721. A rotor is suspended radially by hydrodynamic bearings.In certain embodiments, a radially centered thrust bearing element isprovided to stabilise rotation of the suspended rotor.

U.S. Pat. No. 5,290,227 (Pasque) granted on May 1^(st), 1994 proposes anaxial-flow blood pump having a rotor assembly described generally as ahollowed-out cylinder provided with rotor vanes which extend from theinner surface of the hollowed cylindrical rotor towards the centralrotation axis of the rotor. This design generates two pumping zonesinside the pump, one of these zones being an outer annulus which isexpected to create substantial shearing of the blood in the outer partof the rotor.

European Patent No. EP 0 060 569 granted to Olson et al. on Sep.22^(nd), 1982 teaches a magnetically suspended and rotated impellerwhich comprises a bulky valve member which may be included as part ofthe impeller. Moreover, this European patent teaches impeller bladeswhich axially extend outside of a shroud to connect to the valve member.

Pumps with impeller blades attached to a hollow cylindrical shaft, theshaft rotating around a fixed axle when a magnetic field is applied,reveal a secondary fluid flow path in the same direction as the primarypath. A primary annular flow path is formed between the hollow shaft andthe pump housing, and a secondary annular flow is formed between thehollow shaft and the supporting axle. However, these secondary annularpaths have minimal flow rates and are used to insure proper lubricationof bearing faces or the cleaning of regions which would otherwisecollect debris.

SUMMARY OF THE INVENTION

The present invention relates to a mixed-flow blood pump presentingfeatures of both axial-flow and radial-flow pumps. This mixed-flow bloodpump comprises a stationary housing structure and a rotative impeller.The stationary housing structure defines at least one blood inlet, ablood outlet, and a blood conduit between these blood inlet and bloodoutlet, and the rotative impeller is mounted in the blood conduit. Theat least one blood inlet, the blood outlet, the blood conduit and therotative impeller have respective structures and configurations thatoperate the mixed-flow blood pump at a given point of a maximumhydraulic efficiency curve relating a specific pump rotational speed anda specific pump diameter, that given point being located within atransition region of the maximum hydraulic efficiency curve betweenaxial-flow and radial-flow pumps.

Operating a mixed-flow blood pump, which presents the flowcharacteristics of an axial-flow blood pump while maintaining thethroughput of a radial-flow blood pump, at the given point of themaximum hydraulic efficiency curve located within the transition regionof the curve between radial-flow and axial-flow pumps enables obtentionof a pumping efficiency as high as 53%.

The foregoing and other objects, advantages and features of the presentinvention will become more apparent upon reading of the followingnon-restrictive description of illustrative embodiments thereof, givenby way of example only with reference to the accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

In the appended drawings:

FIG. 1 is a cross sectional view of a human heart in which anillustrative intra-ventricular embodiment of the mixed-flow blood pumpaccording to the present invention is installed;

FIG. 2 is a graph showing, for different types of pumps, a curverelating a specific pump rotation speed N_(s) with a specific pumpdiameter D_(s) at the points where the pump is operating at maximumhydaulic efficiency;

FIG. 3 is a side elevational view of the external shape of theillustrative intra-ventricular embodiment of the mixed-flow blood pumpaccording to the present invention;

FIG. 4 a is a side elevational and cross sectional view of theillustrative intra-ventricular embodiment of the mixed-flow blood pumpaccording to the present invention;

FIG. 4 b is a side elevational and cross sectional view of anillustrative extra-ventricular embodiment of the mixed-flow blood pumpaccording to the present invention;

FIG. 5 a is an enlarged, partial side cross sectional view showing afirst end mount of the drive shaft of the mixed-flow blood pump of FIG.4 a or 4 b;

FIG. 5 b is an enlarged, partial side cross sectional view showing asecond end mount of the drive shaft of the mixed-flow blood pump of FIG.4 a or 4 b; and

FIG. 6 is a schematic view of an illustrative embodiment of a VAD systemimplanted in a human being and comprising the mixed-flow blood pump ofFIG. 4 a.

DESCRIPTION OF THE ILLUSTRATIVE EMBODIMENT

In the following description, important aspects of the pump design areaddressed. In particular, the pump design takes into considerationanatomical and physiological considerations combined with mechanical,electrical and material requirements. Finally, following thisdiscussion, global characteristics of a VAD system are presented.

It should first be noted that the mixed-flow blood pump of the presentinvention is not restricted to an application to an implantable VADsystem. Since the mixed-flow blood pump according to the inventionovercomes a number of drawbacks of the current blood pumps, those ofordinary skill in the art will understand that such a mixed-flow bloodpump can be used as part of an intra-corporal system such as anintra-ventricular VAD, or an extra-ventricular VAD (for example a VADlocated in the abdomen or thorax), or alternatively as a para-corporalor extra-corporal VAD (for example in a bridge to hearttransplantation). It shall also be understood that the mixed-flow bloodpump of the present invention can be used in temporary VADs (for examplea bridge to heart transplant) or permanent VADs.

Anatomical, Physiological and Surgical Considerations

As previously discussed, bleeding is an important problem associatedwith patients who receive a VAD; in fact 30% of patients suffer fromthis problem (Defraigne, J. O., Limet, R., 1996a, “Les assistancescirculatoires: Partie I. Indications et description des systèmes”, RevMed Liege, Vol. 51, pp.295-306). The risk of infection is another quiteimportant problem. These medical and surgical considerations are met bythe illustrative intra-ventricular embodiment of the mixed-flow bloodpump as shown in FIG. 1. This position eliminates the need for inflowand outflow grafts and their anastomoses to thereby reduce the risk ofbleeding and infection. This has also the obvious advantage ofconsiderably reducing the implantation time. FIG. 1 illustrates theproposed position of the illustrative intra-ventricular embodiment ofmixed-flow blood pump 2 in the left ventricle 4.

The mixed-flow blood pump 2 has also been designed to fit in smalladults and in teens. Since the physical size and shape of the mixed-flowblood pump 2 are greatly influenced by the desired location of the pump,a good description of the ventricle anatomy is required. Feigenbaum,Harvey, “Echocardiography”, 5th edition, 1994, Lea & Febiger,Philadelphia, presents several dimensions of the heart normalised by theBSA (Body Surface Area). These anatomical dimensions have beenstatistically determined and are known to represent 95% of thepopulation. Taking into consideration the above-identified statistics, aventricular dimension for humans corresponding to a BSA of 1.5 m² wasused to design the pump.

It will be understood that the physical size and shape of the mixed-flowblood pump 2 could also be adapted to meet the anatomical dimensions ofindividuals falling outside this 95% of the population. Similarly, thesize and shape could be adapted to specific and particular individualsand heart conditions.

For 95% of the population, the internal diameter of the left ventricle 4ranges from 37 to 46 mm in diastole and between 22 to 31 mm in systole.This diameter is determined at the centre of the ventricular length(segment AB in FIG. 1). The diameter near the apex at the first third ofthe ventricular length is about 1.5 cm (segment CD of FIG. 1). Theinternal length of the ventricle from the apex to the aortic valveranges from 55 to 70 mm. Finally, the other important parameter is thesurface of the aortic valve opening which ranges from2.5 to 4 cm².

From a surgical point of view, the favoured insertion. procedure is touse the same approach as with valve replacement. According to thisprocedure, an incision is made at the root of the aorta 6 (FIG. 1) andthe mixed-flow blood pump 2 is inserted though the aortic valve and theninto the left ventricle 4. The mixed-flow blood pump 2 is then pusheduntil its base reaches the myocardium at the apex 8. In order to preventmotion thereof, the mixed-flow blood pump 2 should be fixed. Also inaccordance with an illustrative embodiment, an outflow cannula shouldpass through the aortic valve to further reduce bleeding.

One of the main roles of the mixed-flow blood pump 2 is to restore ahemodynamic function in patients with cardiac failure. Depending on theseverity of the failure and the BSA, the pump 2 is susceptible to workat flow rates between 2 to 6 litres per minute (l/min) against apressure as high as 120 mmHg and, more commonly, at a flow rate between3 to 5 l/min against a pressure of 80 mmHg.

Another important consideration for blood pump design is the hemolysisrate. Hemolysis is the tearing of red blood cells which empties thecontent of the cells in the blood stream resulting in free haemoglobin;the normal level of plasma free haemoglobin is around 10 mg/dl. A bloodpump with a normalised index of hemolysis (NIH) of 0.005 g/100 l andlower is considered to be almost athromatic for red blood cells. A NIHof about 0.05 g/l 100 l could be tolerated. A NIH situated between 0.005g/100 l to 0.05 g/100 l can therefore be envisaged for a VAD. Of course,a NIH as close to 0.005 g/100 l as possible is desirable.

Platelets are other important blood elements; their activation by highhydromechanical forces should be avoided in order to prevent clotformation.

Mixed-flow Flow Blood Pump Design: Mechanical Aspects

This section of the disclosure is divided into parts A, B and C. Part Adescribes the general approach used for the selection of the pumpconfiguration. Part B describes the external shape and size of anillustrative embodiment of the mixed-flow blood pump according to theinvention, and part C describes the internal structure of thisillustrative embodiment of mixed-flow blood pump.

Part A: Selection of a General Pump Configuration

There are three existing non-pulsatile pump configurations, allturbines, and having characteristics which make them potentialcandidates for a cardiac blood pump: radial-flow, axial-flow andmixed-flow pumps. Given their relatively small diameter, cylindricalshape and high throughput, axial-flow pumps display a number ofcharacteristics which make them particularly well suited forimplantation. However, other pump configurations also exhibitcharacteristics, different from those of the axial-flow pump, whichwould also be useful in a cardiac blood pump.

When designing turbine pumps, dimensionless characteristic values areused to compare different pump configurations. Dimensionlesscharacteristic values provide useful indications to pump designers ofexpected performance regardless of the size of the pump, a comparisonwhich would otherwise prove difficult given a virtually infinite numberof operating parameters that depend on infinite variations of internalpump geometry. These dimensionless characteristic values, therefore, canbe used to provide an objective starting point for the selection of ageneral pump configuration.

Two of these dimensionless characteristic values are the specificrotation speed N_(s) of the pump and the specific pump diameter D_(s).They are defined as follows: $\begin{matrix}{N_{s} = \frac{\Omega\quad Q^{1/2}}{H^{3/4}}} & (1) \\{D_{s} = \frac{D \cdot H^{1/4}}{Q^{1/2}}} & (2)\end{matrix}$where Ω is the rotation speed the pump in rad/s, Q is the flow rate inm³/second, H is the head (i.e. the gain in pressure) of the pump and Dthe diameter of the pump, both in meters. N_(s) remains the sameregardless of the size of the pump and therefore provides an accuratemeasure of the performance of a given pump design. D_(s) relates thepump diameter to the pump head H and flow rate Q.

Referring to FIG. 2, known are curves which relate the specific speedN_(s) with the specific diameter D_(s) at the points where the pump isoperating at maximum hydraulic efficiency. In this regard, hydraulicefficiency is expressed as the percentage of the power input to the pumpwhich is converted to energy of movement of the fluid within the pump.From the curve of FIG. 2 and equation (1) above, it follows thatoptimally efficient pumps having a higher specific speed also have asmaller size.

As referred to above, there are three (3) principal categories ofnon-pulsatile pumps characterised by the direction of flow of fluidthrough the pump relative to the axis of rotation: axial-flow,radial-flow and mixed-flow pumps.

In axial-flow pumps the direction of fluid flow is parallel to the axisof rotation. The pressure differential, or head, is produced by a changein the amount of tangential movement. Characteristics associated withaxial-flow pumps include high flow rate Q and small head H. This resultsin high specific speeds N_(s).

In radial-flow pumps a large portion of the throughput, either on theoutlet or the inlet, is radial, i.e. perpendicular to the axis ofrotation. This change in direction causes an increase in pressure.Contrary to an axial-flow pump, a radial-flow pump is characterised byrelatively large head H and smaller flow rate Q, resulting in lowerspecific speeds N_(s).

Located between radial-flow pumps and axial-flow pumps are mixed-flowpumps where the direction of flow at the output or input is composed ofboth radial-flow and axial-flow components. As would be expected, thespecific speed of mixed-flow pumps is located between the specific speedof axial-flow and radial-flow pumps.

In order to determine an optimised choice for a pump, it is necessary toevaluate the specific speed N_(s) in light of the characteristics interms of head H and flow rate Q projected for the pump. As discussedabove, the pump will typically be operated with a flow rate of 5 l/minand a head of approximately 100 mmHg. Additionally, current motortechnology provides small yet efficient motors operating at a speed of11,000 RPM. This gives a specific speed of 1.62.

Referring again to FIG. 2, an indication is given to the ranges of N_(s)and D_(s) within which a given pump configuration will provide efficientoperation. The specific speed N_(s) of 1.62 falls within a transitionregion of the curve between axial-flow and radial-flow pumps. In thistransition region, a mixed-flow pump topology would yield a higherefficiency than purely radial-flow or axial-flow pumps. Additionally,the specific diameter D_(s) is around 2 which, by applying equation (2)above, yields a characteristic diameter of 9.83×10⁻³ for the pump, i.e.a very small pump.

Part B: External Design Requirements

The external design (shape and size) of the mixed-flow blood pump 2(FIG. 1) depends on the anatomic dimensions of the left ventricle 4.FIG. 3 illustrates the external shape of the mixed-flow blood pump 2 andthe critical geometric parameters thereof.

Pump Inlets

FIG. 1 shows that the illustrative embodiment of the mixed-flow bloodpump 2 rests at the bottom of the left ventricle 4, in the region of theapex 8 of the heart 30.

Referring to FIG. 3, in order to prevent the inner walls of the leftventricle 4 from completely obstructing blood intake, two axially spacedapart inlets 12 and 14 are provided. Additionally, a first end of thepump presents the generaly configuration of a hemisphere 16. Thediameter of the hemisphere 16 is set to approximately 20 mm, which issmaller than the segment CD (see FIG. 1) and suitable to limit the levelof pressure on the walls of the left ventricle 4 near the apex 8.

The inlet 12 is found at the first end of the pump 2 between thehemisphere 16 and a first end of an axial, cylindrical member 18. Asillustrated in FIG. 4, the cylindrical member 18 contains the statorwindings 80. A first series of narrow, axially extending supports 20spread out evenly around the axis of the pump 2 connects the hemisphere16 to the cylindrical member 18.

The second inlet 14 is formed between the second end of the cylindricalmember 18 and an impeller housing 22. A second series of narrow, axiallyextending supports 24 spread out evenly around the axis of the pump 2interconnects the cylindrical member 18 to the impeller housing 22. Ascan be seen, the second inlet 14 is axially spaced apart from the firstinlet 12. The separation between the inlets 12 and 14 reduces the effectocclusion of one of the pump inlets may have on normal operation of thepump 2.

Outflow Cannula

Referring back to FIGS. 1 and 3, at the second end of the mixed-flowblood pump 2, the outflow diameter 26 (see FIG. 3) is reduced so as toreduce as much as possible the obstruction caused by an outflow cannula28 to the operation of the aortic valve (not shown); since the functionof the mixed-flow blood pump 2 is to assist blood circulation, bloodflow contribution from the natural contraction of the heart 30 should bemaintained. In an illustrative embodiment, the area of the outflowcannula 28, corresponding to diameter 26, is 1.3 cm².

As illustrated, the outflow cannula 28 is, in the illustrativeembodiment, integral with the impeller housing 22.

As can be better seen from FIG. 3, a blood diffuser 32 is formed at thefree end of the cannula 28, integrally therewith. The function of thediffuser 32 is to reduce the shear stress on blood cells. Withoutdiffuser 32, the velocity of blood ejected from the pump 2 is higherthan the velocity of blood ejected from the heart 30. The difference invelocity between these two blood flows would result in shear stressproportional to this difference. Since the velocity is inverselyproportional to the cross-sectional area, a solution for reducing therelative velocity of the blood flows from the pump 2 and from the heart30 is (a) to increase the area of the orifice 34 of the cannula 28 toreduce the velocity of the flow of blood from the pump 2, and (b) todecrease the area occupied by the blood flow from the heart to increasethe velocity of the latter blood flow. This is exactly the role of thediffuser 32. Of course, parameters such as the angle of opening and thelength of the diffuser 32 can be adjusted at will to fit the mechanicalcharacteristics of the pump 2 in view of minimising the shear stress onthe blood cells.

Housing

The diameter of the mixed-flow blood pump 2 is a compromise betweenpumping requirements and minimal interference with heart contraction. Inan illustrative embodiment, the maximum allowable diameter 36 is about22 mm which is the diameter of the left ventricle 4 in systole. Thisdimension is reasonable since people with heart failure generally havedilated ventricles.

The maximum length 38 of the mixed-flow blood pump 2, as illustrated inFIG. 2, is set in regard of the average distance between the apex 8 andthe aortic valve 40 of the heart 30. In an illustrative embodiment, thelength 38 of the mixed-flow blood pump 2 is about 55 mm. As shown inFIG. 1, a reduction of the pump diameter (see 40) toward the outflowincreases the aortic valve clearance in order to minimise interferencewith the aortic valve function.

Since in this illustrative embodiment the mixed-flow blood pump 2 willbe completely located inside the left ventricle 4, blood will circulatearound the pump 2. As a consequence, all external surfaces of the pump 2should be as smooth as possible and avoid as much as possible abruptdeviations to thereby minimise vortices, turbulence and recirculationzones which may be at the origin of clot formation. To overcome thisproblem, the pump 2 and other components may be machined from surgicalquality titanium.

Fixation Mechanism

At the first end of the pump 2, a fixation mechanism 42 is provided. Asan example, fixation mechanism 42 comprises:

-   -   an elongated needle member 44 extending from the hemisphere 16,        this needle member 44 being driven from the inside of the left        ventricle 4 through the myocardium and the epicardium at the        apex 8 of the heart 30; and    -   a fixation disk 46 fastened to the free end of the needle member        44 on the outside of the heart 30 to firmly fix the mixed-flow        blood pump 2 within the left ventricle 4.

Of course, it is within the scope of the present invention to employ anyother type of fixation mechanism.

Electrical Supply

The required electrical supply for the operation of the motor (to bedescribed herein below) is made through a wire that could, for example,extend from the mixed-flow blood pump 2 along the needle member 44 toreach a controller and an energy source (both to be further describedhereinafter).

Part C: Internal Design Characteristics

In the design of the internal components, some of the characteristics ofaxial-flow pumps have been retained and combined with some of thecharacteristics of radial-flow pumps to form the mixed-flow blood pump2.

Referring to FIG. 4 a, the mixed-flow blood pump 2 comprises astationary housing structure comprising an inflow bushing mount 56 whichdefines the hemisphere 16, the cylindrical member 18, the impellerhousing 22, a stationary outflow stator 52, an outflow bushing mount 58,the outflow cannula 28 and the blood diffuser 32. Blood pump 2 furthercomprises a rotative impeller 50 with an impeller shaft 48 and animpeller blade 70.

Current wet motor axial-flow blood pumps use permanent magnets insertedeither in the central hub or in the impeller blades. Both methodsrequire important compromises. Insertion of the permanent magnets in thecentral hub requires a large hub to locate the permanent magnets closeto the motor windings for obvious electromagnetic coupling reasons. Incontrast, a small hub has the advantage of increasing the pumped volumeof blood. Insertion of the permanent magnets in the impeller bladesyields a compromise since the geometry of the blades must be curved forpumping efficiency. As a consequence, some of the embedded magnets moveaway from the windings as the blade(s) curve(s).

In the approach proposed by the illustrative embodiment of the presentinvention, the mixed-flow blood pump 2 presents an enclosed-impellermixed-flow configuration. In the illustrative intra-ventricularembodiment of this configuration as illustrated in FIG. 4 a, theimpeller blade 70 is a spiralling auger-type impeller blade having aconstant height and being rigidly attached to the outer surface of theimpeller drive shaft 48 and enclosed in the impeller housing 22.Obviously, the impeller blade 70 fits snugly withing the impellerhousing 22. Of course, the shape (curvature and angulation) of theimpeller blade 70 should be optimally designed in relation to pumpingperformance and other hydrodynamic considerations. In particular, theinfluence of the blade angulation on the level of shearing stresses,turbulence and cavitation responsible for red blood cell damage andincrease of hemolysis rate must be carefully taken into consideration.

The end portion of the impeller shaft 48 bearing the impeller blade 70is slightly tapered in a direction opposite to the direction of bloodflow. This contributes to create the mixed-flow operation of the pump 2.More specifically, this slight taper imparts to the blood flow bothaxial and radial components.

Still referring to FIG. 4 a, at both ends of the impeller drive shaft 48end pivots 66 and 68 protrude. The function of the end pivots 66 and 68is to support the impeller drive shaft 48 at each end. The end pivots 66and 68 are respectively inserted into the respective bushing mounts 56and 58. Bushing 72 is mounted on the inner face of the inflow bushingmount 56. In the same manner, bushing 74 is mounted on the adjacent faceof the outflow bushing mount 58. The bearings formed by the bushing andpivot assemblies 66;72 and 68;74 are the only mechanical parts subjectto wear. Therefore, these parts are expected to be mostly responsiblefor the life span of the mixed-flow blood pump 2.

Still referring to FIG. 4 a, the cylindrical gap 76 separating theimpeller shaft 48 and the inner surface of the cylindrical member 18should be sufficiently thick to produce sufficient blood flow in orderto increase washout and prevent clot formation. On the other hand, toolarge a gap 76 may either reduce the pump efficiency (by reducing theelectromagnetic coupling) or result in higher hemolysis. Just a word tomention that the section of the impeller shaft 48 within the cylindricalmember 18 is cylindrical whereby the gap 76 has a constant thickness.

In the illustrative embodiment of FIGS. 1 and 4 a, the volume of bloodpumped through the second inlet 14 is typically 4 liters/minute. This ishigher than the volume of blood pumped through the first inlet 12 andthe cylindrical gap 76 which is typically 1 liter/minute. A number ofbenefits is associated with the higher volume blood pumped through thesecond inlet 14. For example, installation of the mixed-flow blood pump2 in the left ventricle 4 of a patient with the cannula 28 extendingthrough the aortic valve generally interferes with proper operation ofthe aortic valve. Optimally, the aortic valve should continue tofunction normally; however, in some cases, it has been observed that theaortic valve ceases to function further until it remains closed aroundthe cannula 28. Typically, blood would have the tendency to collect inthe region close to the aortic valve and the cannula 28 which might leadto thrombus formation and other adverse effects. The increased volume ofblood pumped through the second inlet 14 has the effect of creatingturbulence in the region within the ventricle 4 bordered by the aorticvalve and the cannula 28, thus providing improved washout of this regionand thereby reducing the effects of the malfunctioning aortic valve.

On the one hand, the volume of blood pumped through the second inlet 14contributes to the radial-flow operation of the mixed-flow blood pump 2.On the other hand, the volume of blood pumped through the first inlet 12and the cylindrical gap 76 contributes to the axial-flow operation ofthe mixed-flow blood pump 2.

The stationary outflow stator 52 comprises a plurality of blades shapedand disposed around the outflow bushing mount 58 to transform therotational motion of the flow about the longitudinal axis 54 into atranslational motion. Therefore, the stationary outflow stator 52constitutes a blood flow straightener.

As previously mentioned, the pump design should minimise shearing stressin order to minimise hemolysis. In that context, reduction of therotational speed would obviously contribute to reduce hemolysis.However, reduction of the rotational speed while pumping the same volumeof blood requires an increase of the volume of blood contained in therotor zone of the pump 2. The volume of blood contained in the rotorzone can be increased by either increasing the diameter of the rotorzone, or alternatively minimising the volume of the central hub of thepump rotor.

Design of the internal surfaces of the pump is also important forminimising shearing stresses in order to minimise hemolysis. FIG. 5 ashows the configuration of the region between the inflow bushing mount56 and the impeller drive shaft 48. FIG. 5 b shows the configuration ofthe region between the impeller drive shaft 48 and the outflow bushingmount 58.

Referring to FIG. 5 a, the end surface 60 of the impeller drive shaft 48is convex and generally hemispheric while the confronting surface 62 ofthe bushing mount 56 is concave and generally hemispheric. Forming thesurfaces in this manner reduces the shearing stress placed on the bloodthus minimising hemolysis (see flow 61).

The outflow bushing mount 58 is formed with a similar convex, generallyhemispheric end surface to reduce shearing stress at the outlet (seeFIG. 5 b, in particular flow 65). The confronting end of the impellershaft 48 is flat and perpendicular to the axis 54.

Additionally, referring to both FIGS. 5 a and 5 b, end pivots 66 and 68by which the respective ends of the impeller shaft 48 are mounted totheir respective bushing mounts 56, 58 are slightly tapered in adirection opposite to blood flow. This helps prevent the creation ofeddies and the collection of debris in proximity to the end pivots.Normally, a taper of the order of five degrees (5°) is adequate.

In this manner, pivot 66 has a smaller diameter free end received withinthe bushing 72. In the same manner, the pivot 68 has a larger diameterfree end received in the bushing 74.

FIG. 4 b illustrates an alternative, illustrative extra-ventricularembodiment of mixed-flow blood pump 100. Pump 100 is adapted for useexternally of the heart as a ventricle bypass/assist. In this embodimentthe pump 100 would typically be implanted above the diaphragm in thethorax and would be connected to the circulation system using standardvascular grafts, a first graft 102 being attached to the inflow end ofthe pump and a second graft 104 being attached to the outflow end of thepump 100.

Similar to the mixed-flow blood pump 2, the alternative illustrativeembodiment 100 as illustrated in FIG. 4 b comprises a stationary housingstructure including an inflow bushing mount 122, a cylindrical member108, an impeller housing 106, a stationary outflow stator 114, anoutflow bushing mount 124, an outflow cannula 150 and a blood diffuser151. Blood pump 100 further comprises a rotative impeller 112 with animpeller shaft 110 and an impeller blade 116.

The impeller blade 116 is a spiralling auger-type impeller blade rigidlyconnected to the impeller drive shaft 110 and enclosed in the impellerhousing 106.

A series of longitudinal ridges 118, typically five (5) evenly spacedaround the pump axis 119, support the cylindrical member 108 within theimpeller housing 106 thereby forming a series of longitudinal flowpassages such as 120 between the cylindrical member 108 and the impellerhousing 106.

The longitudinal ridges 118 extend to meet and hold rigid the inflowbushing mount 122 with respect to the impeller housing 106 and thecylindrical member 108. Similarly, the outflow bushing mount 124 issupported within the impeller housing 106 through the stationary outflowstator 114.

The stationary outflow stator 114 comprises a plurality of blades shapedand disposed around the outflow bushing mount 124 to transform therotational motion of the blood flow about the longitudinal axis 119 intoa translational motion. Therefore, the stationary outflow stator 114constitutes a flow straightener.

At both ends of the impeller shaft 110 end pivots 126 and 128 protrude.The end pivots 126 and 128 are respectively inserted into respectivebushings 130 and 132 to support the impeller drive shaft 110 within thecylindrical conduit 108 while at the same time allowing the impellershaft 110 to rotate freely. Bushing 130 is mounted on the inflow bushingmount 122. In the same manner, bushing 132 is mounted on the outflowbushing mount 124.

In addition to the series of longitudinal flow passages 120, an annularflow passage 134 is formed between the inner surface 136 of thecylindrical member 108 and the outer surface 138 of the impeller shaft110.

Flow of blood through the mixed-flow pump 100 is indicated by the arrows152-154.

The other characteristics of the mixed-flow pump 2 according to theillustrative embodiment of FIG. 4 a also apply to the illustrativeembodiment 100 of FIG. 4 b.

Electrical Aspects

In the illustrative embodiment of FIG. 4 a, the mixed-flow blood pump 2is actuated by means of a brushless DC (direct current) motor. Thisbrushless configuration presents the advantage of minimal wear. Twoother interesting characteristics of brushless DC motors are highrotational speed and high torque.

In the mixed-flow blood pump 2, the brushless DC motor includeselongated axial permanent magnets such as 78 inserted in the impellershaft 48 and stator windings 80 embedded or housed in the cylindricalmember 18.

As discussed previously, the cylindrical gap 76 between the outersurface of the impeller shaft 48 and the inner surface of thecylindrical member 18 must be sufficiently thick to produce sufficientblood flow in order to increase washout and prevent clot formation.However, increasing the thickness of the gap 76 decreases the efficiencyof the magnetic coupling between the permanent magnets 78 and the statorwindings 80. This requires an increase in current through the statorwindings 80 to compensate for the decreased efficiency and to maintainthe same characteristics in terms of impeller blade speed and bloodvolume throughput. Of course, increase in current leads to an increasein thermal loss from the stator windings 80; this thermal loss increasesas the square of the current through the stator windings 80. As thetemperature of the surface of the stator windings must remain at orbelow 40° C., the gap 76 must be sufficiently small to provide efficientmagnetic coupling between the permanent magnets 78 and the statorwindings 80.

Thermal performance is also improved given the proximate position of thestator windings 80 to the external surface 82 of the cylindrical member18. Blood flow over the external surface 82 efficiently cools the statorwindings 80. The flow of blood within the gap 76 between the impellershaft 48 and the inner surface 84 of the cylindrical member 18 alsocontributes in efficiently cooling the stator windings 80.

The alternative illustrative embodiment 100 of the mixed-flow blood pumpas illustrated in FIG. 4 b maintains the essential electricalcharacteristics of the illustrative embodiment of FIG. 4 a with theexception that, referring to FIG. 4 b, the design of the pump 100overcomes the thermal limitations by allowing for a second blood flowpassage along the series of longitudinal flow passages 120 between theimpeller housing 106 and the cylindrical member 108.

Axial spacing between the impeller blade and the permanent magnets alongthe impeller shaft enables separate design of the DC motor and theimpeller to obtain simultaneously both efficient coupling between thepermanent magnets and the stator windings and sufficient pumping volume.

Selection of the Materials

The choice of materials for an implantable device is crucial and severalproperties of the available materials should be considered: strength,durability, hardness, elasticity, wear resistance, surface finish andbiocompatibility. Biocompatibility is very important to minimiseirritation, rejection and thrombogenesis. The interaction between thesurface of the material and the biological tissues is very complex. Inseveral cases, treatment of the surface with human proteins, certaindrugs like heparin or other biocompatible material may considerablyincrease the biocompatibility and minimise thrombus formation.

VAD System

FIG. 6 schematically illustrates an embodiment of implantable VAD systemincluding an axial-flow blood pump 2. The VAD system is composed of fourmain parts:

-   -   the axial-flow blood pump 2 implanted in the left ventricle 4 of        the patient 86;    -   an internal controller 88;    -   two energy sources, namely an internal rechargeable battery 90        and an external rechargeable battery 92; and    -   a Transcutaneous Energy and Information Transmission (TEIT)        system 94.

VAD and TEIT Systems are well known in the art and will not be furtherdiscussed in the present specification.

To conclude, ventricular assist devices (VADs) are now being usedworld-wide and their utilisation is becoming more and more accepted as asolution to treat end stage heart failure. It is generally accepted thatVADs extend life of patients while improving quality of life of thesepatients. A poll, made with patients who received VADs, concerning theirquality of life revealed that these patients would have preferred aheart transplant but prefer their situation than having to be ondialyses.

It is also now being accepted that VAD is becoming a cost effectivesolution considering the fact that patients are discharged from thehospitals more rapidly and may return to normal life occupations. In theUnited States, several insurance companies are now reimbursing theimplantation of VADs.

Finally, the mixed-flow blood pump 2 according to the invention providesan excellent bridge to heart transplant and aims at long term implant.The new proposed mixed-flow blood pump 2 should answer most of theremaining problems and limitations of the prior axial-flow blood pumps,especially those related to hemolysis and bleeding.

Although the present invention has been described hereinabove by way ofillustrative embodiments thereof, these embodiments can be modified atwill, within the scope of the appended claims, without departing fromthe spirit and nature of the present invention.

1. A mixed-flow blood pump presenting features of both axial-flow andradial-flow pumps, comprising: a stationary housing structure defining alongitudinal axis, an axially-extending annular blood inlet passage, aradially-extending annular blood inlet passage, an axial blood outlet,and an axial blood conduit between (a) the axially-extending andradially-extending blood inlet passages, and (b) the axial blood outlet;and an rotative impeller mounted within the stationary housingstructure, comprising: an impeller shaft rotative about the longitudinalaxis of the stationary housing structure, the impeller shaft having, inthe axial blood conduit, a shaft portion tapered in a direction oppositeto the direction of blood flow; and an impeller blade mounted on thetapered shaft portion; wherein the axially-extending annular blood inletpassage, the radially-extending annular blood inlet passage, the taperedshaft portion, the impeller blade mounted on the tapered shaft portion,and the axial blood outlet operate the mixed-flow blood pump at a givenpoint of a maximum hydraulic efficiency curve relating a specific pumprotational speed and a specific pump diameter, said given point beinglocated within a transition region of the maximum hydraulic efficiencycurve between axial-flow and radial-flow pumps.
 2. A mixed-flow bloodpump as defined in claim 1, wherein the specific pump rotation speed andthe specific pump diameter have dimensionless values.
 3. A mixed-flowblood pump as defined in claim 2, wherein the dimensionless values ofthe specific pump rotational speed is 1.62.
 4. A mixed-flow blood pumpas defined in claim 1, wherein: the stationary housing structurecomprises first and second axially spaced apart annular blood inlets;the axially-extending annular blood inlet passage extends between thefirst annular blood inlet and the axial blood conduit; and theradially-extending annular blood inlet passage extends between thesecond annular blood inlet and the axial blood conduit.
 5. A mixed-flowblood pump as defined in claim 4, wherein: the axially-extending annularblood inlet passage comprises a generally cylindrical axial passageportion.
 6. A mixed-flow blood pump as defined in claim 4, wherein: theradially-extending annular blood inlet passage defines an acute anglewith respect to the longitudinal axis of the stationary housing.
 7. Amixed-flow blood pump as defined in claim 5, wherein: the stationaryhousing structure comprises a cylindrical member with an inner surface;and the generally cylindrical, axial passage portion comprises a gapbetween the impeller shaft and the inner surface of the cylindricalmember.
 8. A mixed-flow blood pump as defined in claim 1, in which theimpeller blade comprise an auger-type impeller blade mounted on theimpeller shaft.
 9. A mixed-flow blood pump as defined in claim 1,wherein: the stationary housing structure comprises a cylindrical memberaround the impeller shaft; and the mixed-flow blood pump furthercomprises an electrical motor structure comprising: permanent magnetsembedded within the impeller shaft; and electrical windings mountedwithin the cylindrical member of the stationary housing structure.
 10. Amixed-flow blood pump as defined in claim 7, wherein: the mixed-flowblood pump further comprises, in the region of the gap between theimpeller shaft and the inner surface of the cylindrical member, anelectrical motor structure comprising: permanent magnets embedded withinthe impeller shaft; and electrical windings mounted within thecylindrical member of the stationary housing structure.
 11. A mixed-flowblood pump as defined in claim 1, wherein: the impeller shaft comprisesfirst and second opposite ends and first and second opposite end pivots;and the stationary housing structure comprises first and second bushingmounts at the respective first and second ends of the impeller shaft toreceive the first and second opposite end pivots, respectively.
 12. Amixed-flow blood pump as defined in claim 11, wherein: the first pivotis slightly tapered in a direction opposite to the direction of bloodflow and comprises a smaller diameter free end; the first bushing mountcomprises a first bushing to receive the smaller diameter free end ofthe first pivot; the first bushing mount comprises a concave, generallyhemispheric surface in the region of the first bushing; and the firstend of the impeller shaft is convex and generally hemispheric.
 13. Amixed-flow blood pump as defined in claim 11, wherein: the second pivotis slightly tapered in a direction opposite to the direction of bloodflow and comprises a larger diameter free end; the second bushing mountcomprises a second bushing to receive the larger diameter free end ofthe second pivot; the second bushing mount comprises a convex, generallyhemispheric surface in the region of the second bushing; and the secondend of the impeller shaft is generally flat.
 14. A mixed-flow blood pumpas defined in claim 1, wherein: the stationary housing structurecomprises an impeller housing around the impeller blade, the impellerhousing having an inner surface in which the impeller blade snugly fits.15. A mixed flow blood pump as defined in claim 11, wherein: theimpeller blade has a constant height.
 16. A mixed-flow blood pump asdefined in claim 14, further comprising: a flow straightener structureconnected to the impeller housing for straightening the blood flow fromthe impeller blade; and a flow diffuser structure downstream the flowstraightener structure to diffuse the straightened blood flow.